Magnetic resonance imaging (MRI) is a noninvasive imaging modality capable of distinguishing a wide variety of objects based on their intrinsic composition and also is an imaging technique that is capable of providing one-, two- or three-dimensional imaging of the object. A conventional MRI system typically includes a main or primary magnet that provides the main static magnetic field Bo, magnetic field gradient coils and radio frequency (RF) coils, which are used for spatial encoding, exciting and detecting the nuclei for imaging. Typically, the main magnet is designed to provide a homogeneous magnetic field in an internal region within the main magnet, for example, in the air space of a large central bore of a solenoid or in the air gap between the magnetic pole plates of a C-type magnet. The patient or object to be imaged is positioned in the homogeneous field region located in such air space. The gradient field and the RF coils are typically located external to the patient or object to be imaged and inside the geometry of the main or primary magnet(s) surrounding the air space. There is shown in U.S. Pat. Nos. 4,689,563; 4,968,937 and 5,990,681, the teachings of which are incorporated herein by reference, some exemplary MRI systems.
In MRI, the uniform magnetic field Bo generated by the main magnet is applied to an imaged object by convention along the Z-axis of a Cartesian coordinate system, the origin of which is within the imaged object. The uniform magnetic field Bo being applied has the effect of aligning the magnetization arising from the nuclei of the atoms comprising the imaged object, along the Z-axis, such nuclei possess a nuclear magnetization due to their having an odd number of protons or neutrons. In response to RF magnetic field pulses of the proper frequency, with field direction orientated within the XY plane, the nuclei resonate at their Larmor frequencies, ω=γBo where γ is called the gyromagnetic ratio. In a typical planar imaging sequence, the RF signal centered about the desired Larmor frequency is applied to the imaged object at the same time a magnetic field gradient Gz is being applied along the Z-axis. This gradient field Gz causes only the nuclei in a slice of limited thickness through the object perpendicular to the Z-axis, to satisfy the resonant condition and thus be excited into resonance.
After excitation of the nuclei in the slice, magnetic field gradients are applied along the X- and Y-axes respectively. The gradient Gx along the X-axis causes the nuclei to precess at different frequencies depending on their position along the X-axis, that is, Gx spatially encodes the precessing nuclei by frequency. Thus, this gradient is often referred to as a frequency encoding or read-out gradient. The Y-axis gradient Gy is incremented through a series of values and encodes the Y position into the rate of change of the phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
The quality of the image produced by the MRI techniques is dependent, in part, upon the strength of the magnetic resonance (MR) signal received from the precessing nuclei. For this reason an independent RF coil is often placed in close proximity to the region of interest of the imaged object, more particularly on the surface of the imaged object, in order to improve the strength of the received signal. Such RF coils are sometimes referred to as local coils or surface coils.
There is described in U.S. Pat. No. 4,825,162 a surface coil(s) for use in MRI/NMRI imaging and methods related thereto. In the preferred embodiment of that invention, each surface coil is connected to the input of an associated one of a like plurality of low-input-impedance preamplifiers, which minimizes the interaction between any surface coil and any other surface coils not immediately adjacent thereto. These surface coils can have square, circular and the like geometries. This yields an array of a plurality of closely spaced surface coils, each positioned so as to have substantially no interaction with all adjacent surface coils. A different MR response signal is received at each different one of the surface coils from an associated portion of the sample enclosed within the imaging volume defined by the array. Each different MR response signal is used to construct a different one of plurality of different images from each surface coil. These images are then combined, on a point-by-point basis to produce a single composite MR image of a total sample portion comprised of the MR response signals from the entire array of surface coils.
Common MRI strategies use the relative density of water protons in a sample and the relaxation properties of energy exchange between the local spins with the lattice and with each other to achieve contrast between tissues. The spin-lattice interaction is described by the T1 relaxation time, and the spin-spin interaction is described by the T2 relaxation time. Furthermore, static field inhomogeneities directly affect T2 relaxivity as the local field variations introduce rapid intravoxel dephasing. These inhomogeneities can arise from a variety of sources, but significant mismatches in tissue magnetic susceptibility coefficients between adjacent tissues or objects are an important source. As a result, transverse signal decay occurs much faster and the modified relaxation time is referred to as T2*. Thus, MRI pulse sequences are conventionally designed to extract tissue contrast by exploiting intrinsic differences in the proton density, T1, T2, and T2*.
In general, MRI relies heavily on a homogeneous static field for accurate imaging. Additionally, the object or objects being imaged commonly exhibit a diamagnetic behavior, which means that their magnetic susceptibility values (χ) are both small and negative. In contrast, non-diamagnetic objects may have χ values that are small and positive (paramagnetic) or large and positive (ferromagnetic). Objects that create magnetic field susceptibility typically lead to signal voids in MRI that extend well beyond the area encompassed by these objects (i.e., “bloom artifact”). Moreover, there are many potential sources of signal voids (e.g., absence of signal-generating substance, motion artifacts, magnetic field inhomogeneity, etc.) that have similar hypointense signal intensity in images.
The creation of the signal void is related to the magnetic susceptibility gradients (MSGs) created by the non-diamagnetic object. In turn, MSGs are dependent on many factors, including the magnetic susceptibility value of the non-diamagnetic objects, the environment in which the non-diamagnetic object is placed, the type of pulse sequence used, and the individual imaging parameters associated with the pulse sequence. Therefore, the specificity for localization of susceptibility-generating objects is suboptimal.
Since non-diamagnetic objects affect T1 and T2 relaxation properties, imaging strategies have been adopted either to minimize or maximize the T2* effects of these agents depending on whether the goal is to minimize or maximize the sensitivity to the object, respectively. One method is to acquire T1 and/or T2 maps, which is a time-intensive process, lacks sensitivity, and is difficult to acquire on moving structures.
Other common methods for detecting susceptibility-inducing objects are T2*-weighted images using spin echo (SE) [Hahn, E. L. Spin Echoes. Phys Rev 1950; 80 (4): 580-594], fast spin echo (FSE) [Henning, J., Nauerth, A. and H. Friedburg. RARE imaging: a fast imaging method for clinical MR. Magn Reson Med 1986; 3: 823-833], or gradient echo (GRE) [van der Meulen, P., Groen, J. P. and J. J. Cuppen. Very fast MR imaging by field echoes and small angle excitation. Magn Reson Imaging 1985; 3 (3): 297-299] imaging. T2*-weighted imaging is quite sensitive to MSGs but often creates large signal voids that extend far beyond the source object. Thus, the relationship between the non-diamagnetic object volume and the volume of the hypointensity may not be linear. Furthermore, the negative contrast associated with non-diamagnetic objects can be difficult to discriminate from other potential sources of signal voids (i.e., the absence of tissue, motion artifacts, calcifications, hemorrhage, etc.).
Techniques have been recently proposed to create positive contrast from the susceptibility artifacts created by such non-diamagnetic objects as superparamagnetic iron oxide nanoparticles [Cunningham, C. H., Arai, T., Yang. P., M., McConnell, M. V., Pauly, S. M., and S. M. Conolly. (2004). Positive contrast magnetic imaging of cells labeled with magnetic nanoparticles. Magn Reson Med 2005; 53 (5):999-1005 (2005); Coristine, A. S., Foster, P. J., Deoni, S. C., Heyn, C. and B. K. Rutt (2004). Positive contrast labeling of SPIO loaded cells in cell samples and spinal cord injury. Proc Intl Soc Magn Reson Med 11: 163; and Mani, V., Briley-Saebo, K. C., Itskovich, V. V., Samber, D. D. and Z. A. Fayad. Gradient echo acquisition for superparamagnetic particles with positive contrast (GRASP): sequence characterization in membrane and glass superparamagnetic iron oxide phantoms at 1.5 T and 3 T. Magn Reson Med 2006; 55 (1):126-135.] and paramagnetic ring-tipped catheters [Seppenwoolde, J. H., Viergever, M. A. and Bakker, C. J. Passive tracking exploiting local signal conservation: the white marker phenomenon. Magn Reson Med 2003; 50 (4): 784-790]. The technique proposed by Cunningham et al uses spectrally-selective radio frequency pulses to image a spectral band assumed to be associated with the MSGs. The techniques described in Coristine et al., Seppenwoolde et al., and Mani et al. use a “white marker” method where signal enhancement is created using gradient imbalances such that only a set of spectral frequencies associated with the MSGs is refocused. Both of these methods produce a positive signal, rather than signal voids, however, this is achieved at the cost of a large signal loss. Furthermore, the technique by Cunningham et al. is sensitive to the tuning of the passband and may further attenuate signal from the MSGs. The techniques used by Seppenwoolde et al., Mani et al., and Coristine et al. also require careful consideration of imaging parameters.
Also, interventional MRI is playing an increasingly important role for guiding endovascular procedures. MR-compatible devices for these procedures are often designed to be non-magnetic or to be made from magnetic resonance (MR) invisible using materials, such as nitinol. Thus, localization of the device becomes difficult. To make MR-compatible devices MR-visible, either passive markers are incorporated [Bakker, C. J., Hoogeveen, R. M., Weber, J., van Vaals, J. J., Viergever, M. A. and W. P. Mali. Visualization of dedicated catheters using fast scanning techniques with potential for MR-guided vascular interventions. Magn Reson Med 1996; 36 (6): 816-820; Bakker, C. J., Hoogeveen, R. M., Hurtak, W. F., van Vaals, J. J., Viergever, M. A. and W. P. Mali. MR-guided endovascular interventions: susceptibility-based catheter and near-real-time imaging technique. Radiology 1997; 202 (1): 273-276; Bakker, C. J., Bos, C. and H. J. Weinmann. Passive tracking of catheters and guidewires by contrast-enhanced fluoroscopy. Magn Reson Med 2001; 45 (1): 17-23; Wacker, F. K., Reither, K., Branding, G., Wendt, M. and K. J. Wolf. Magnetic resonance-guided vascular catheterization: feasibility using a passive tracking technique at 0.2 Tesla in a pig model. J Magn Reson Imaging 1999; 10 (5): 841-844], or active coil tracking systems are incorporated [Atalar, E., Bottomley, P. A., Ocali, O., Correia, L. C., Kelemen, M. D., Lima, J. A. and E. A. Zerhouni. High resolution intravascular MRI and MRS by using a catheter receiver coil. Magn Reson Med 1996; 36 (4): 596-605; Dumoulin, C. L., Souza, S. P. and R. D. Darrow. Real-time monitoring of invasive devices using magnetic resonance. Magn Reson Med 1993; 29 (3): 411-415; Erhart, P., Ladd, M. E., Steiner, P., Heske, N., Dumoulin, C. L. and J. F. Debatin. Tissue-independent MR tracking of invasive devices with an internal signal source. Magn Reson Med 1998; 39 (2): 279-284; Ladd, M. E., Zimmermann, G. G., McKinnon, G. C., von Schulthess, G. K., Dumoulin, C. L., Darrow, R. D., Hofmann, E. and J. F. Debatin. Visualization of vascular guidewires using MR tracking. J Magn Reson Imaging 1998; 8 (1): 251-253; Wendt, M., Busch, M., Wetzler, R., Zhang, Q., Melzer, A., Wacker, F., Duerk, J. L. and J. S. Lewin. Shifted rotated keyhole imaging and active tip-tracking for interventional procedure guidance. J Magn Reson Imaging 1998; 8 (1): 258-261] are incorporated. Such incorporation with the device, however, has the negative impact of increasing the device diameter and device complexity.
In addition, chemical analysis of iron from liver needle biopsy specimens is currently the most accepted method of diagnosis of iron storage diseases such as hemochromatosis and thalassemia. However, hepatic iron concentration (HIC) measurements by needle biopsy have several problems including: sampling errors owing to the large variation in HIC from site to site within the liver and the uncomfortable nature of the procedure for patients which limits testing frequency. MRI methods have been used to non-invasively assess iron overload in the liver and heart [Anderson, L. J., Holden, S., Davis, B., Prescott, E. Charrier, C. C., Bunce, N. H., Firmin, D. N., Wonke, B., Porter, J., Walker, J. M. and D. J. Pennell. Cardiovascular T2-star (T2*) magnetic resonance for the early diagnosis of myocardial iron overload. Eur Heart J 2001; 22: 2171-2179; Clark, P. R. and T. G. St. Pierre. Quantitative 1/T2 mapping of hepatic iron overload: a single spin echo imaging methodology. Magn Reson Imaging 2000; 18: 431-438; Jensen, P. D., Jensen, F. T., Christensen, T., Eiskjaer, H., Baandrup, U. and J. L. Nielsen. Evaluation of myocardial iron by magnetic resonance imaging during iron chelation therapy with deferrioxamine: indication of close relation between myocardial iron content and chelatable iron pool. Blood 2003; 101: 4632-4639; Mavrogeni, S. I. Gotsis, E. D., Markussis, V., Tsekos, N., Politis, C., Vretou, E. and D. Kermastinos. T2 relaxation time study of iron overload in β-thalassemia. MAGMA 1998; 6 (1): 7-12; Westwood, M., Anderson, L. J., Firmin, D. J., Gatehouse, P. D., Charrier, C. C. Wonke, B. and D. J. Pennell. A single breath-hold multiecho T2* cardiovascular magnetic resonance technique for diagnosis of myocardial iron overload. J Magn Reson Imaging 2003; 18: 33-39], thereby overcoming sampling errors and allowing more frequent testing. There is still a need, however, for non-invasive measurement techniques that are more sensitive for iron measurement.
A non-invasive MRI technique commonly referred to as BOLD MRI [Ogawa, S., Lee, T. M., Kay, A. R. and D. W. Tank. Brain magnetic resonance imaging with contrast dependent on blood oxygenation. Proc Natl Acad Sci USA 1990; 87 (24): 9868-9872] has been used to assess blood oxygenation, where differences in blood oxygenation are used to modulate signal intensity. Briefly, hemoglobin is the iron-containing respiratory protein of red blood cells that transports oxygen as oxyhemoglobin from the lungs to the tissues. Following delivery of oxygen to the tissues, the oxyhemoglobin becomes deoxyhemoglobin with a resulting change in the magnetic properties of the blood. Deoxyhemoglobin is paramagnetic and thus produces intravascular bulk magnetic field gradients in and around the surrounding tissue. BOLD MRI uses a subtraction technique with T2*-weighted imaging to study blood flow and oxygen utilization. Because of the small signal change generated by the BOLD effect, however, BOLD MRI is typically performed on high-field scanners (3 T or greater). Thus, there continues to be a need to assess blood oxygenation using MRI scanners at lower field strengths (e.g., 1.5 T).
The references discussed herein are provided solely for their disclosure prior to the filing date of the present application. Nothing herein is to be construed as an admission that the inventors are not entitled to antedate such disclosure by virtue of prior invention.
It thus would be desirable to provide a methodology for acquisition of MRI image data where susceptibility-generating objects, including preferably other sources of static field inhomogeneities, as hyperintense signals with the geometric extent of the signal-enhancing effect being controlled by imaging parameters that can be varied as well as MRI systems, MRI apparatuses and software applications programs embodying such methods. It would be particularly desirable to provide such systems, methods and applications programs that would embody a positive contrast method that uses selective suppression methods to attenuate some signals and enhancing signals associated with magnetic susceptibility gradients. It would be particularly desirable to provide such systems, methods and applications programs that would attenuate signals from fat and from water, more specifically on-resonant water. It would be particularly desirable to provide such systems, methods and applications programs that would attenuate signals such as those for fat and/or on-resonant water protons and enhance signals associated with magnetic susceptibility gradients, thereby improving contrast-to-noise ratio and specificity of MSGs in comparison to prior art methods, systems and/or applications programs.